For over a hundred years photographic films have been used to capture and display x-rays for diagnostic purposes. In recent years, digital radiography (DR) has become increasingly popular. DR refers to the application of digital equipment and image processing techniques to projection radiography. Digitally recorded x-rays are superior to those recorded with photographic film due to the greater dynamic range offered by a digital recording system. Furthermore, computer image processing techniques provide a wealth of capabilities to study otherwise obscured details within the image.
One type of DR imaging device is an optically-coupled charge-coupled device (CCD) DR system used for clinical diagnosis. Typical optically coupled CCD-based DR systems use a scintillator screen, a mirror and a lens to capture and reduce an x-ray image onto a CCD camera for digitization. To take a digital radiograph using such a system, a DR imaging unit is positioned behind a subject. A standard radiographic generator positioned in front of the subject directs radiation through the subject to a fluorescent-imaging scintillator screen mounted just behind the front surface of the imaging unit. The scintillator screen is the conversion media for radiation to visible light. The scintillator screen absorbs the radiographic radiation and emits light of a particular wavelength which closely matches the peak sensitivity of a CCD camera. A front-surfaced mirror is positioned at an angle inside the imaging unit to direct the visible radiographic image into the CCD camera. The mirror allows the CCD camera to be positioned out of the direct path of the radiation, effectively shielding it from radiation exposure and prolonging its life. A high-efficiency lens is located between the mirror and camera and reduces the image and directs it onto the surface of a CCD sensor in the camera.
The visual image formed by the fluorescent-imaging screen is converted into a digital image by the CCD sensor. A control computer converts the image into a medical image file that can be viewed for clinical diagnosis, enhanced and electronically stored with patient demographic information in a picture archiving system.
Digital radiography has enabled the use of a technique known as dual energy subtraction radiography, which exploits the energy dependence of x-ray attenuation by different tissues. When producing multiple images of a subject obtained by multiple x-ray exposures at different kilovolt peak (kVp) levels and/or by a different filtering of a single x-ray exposure, the photons will interact differently in the scintillator and/or subject. The proportion of photoelectric absorption to Compton scattering will be different in the generation of the different images. Using this effect, a third image can be calculated from the two, in which for instance, the bone structure or soft tissue can be significantly enhanced or suppressed.
One known approach to dual energy digital imaging involves digital imaging devices that use sequential x-ray exposures in rapid succession, at different kVp settings. A scintillator produces multiple images when struck by the multiple x-ray exposures, and these images are captured by a digital sensor for image processing. Because this technique involves multiple sequential exposures, the time delay between exposures tends to cause misregistration resulting in a less-than-perfect separation of the bone and soft tissue components.
Another application of this technique uses a single x-ray exposure detected by two detectors separated by a filter. The filter attenuates a portion of the x-ray spectrum, thereby enabling the detectors to produce two images of the same subject but with different kVp levels, and different contrast properties. Using these two images will make it possible, for instance, to separate the bone structures in one image from the other image, thereby generating a third image that primarily shows soft tissue. Examples of such applications are disclosed in U.S. Pat. No. 4,626,688 (Barnes) and CA 2,218,127 and CA 2,254,877 (Karellas). In Barnes, the detectors are photodiodes that are both located directly in the path of the x-ray exposure, and in Karellas, the detectors are CCD detectors, of which at least one of the detectors are located in the x-ray exposure, along with other components. Disadvantageously, the systems disclosed in both Barnes and Karellas both locate numerous components other than a filter in the path of the x-ray exposure between the x-ray source and the detector. Such non-filter components can corrupt the x-ray image recorded by the detector. Also, both systems locate at least one detector in the direct path of the x-ray exposure, which can be harmful to the detector.